Many types of digital imaging devices create image data using stored electrical charge. For example, known charge-coupled devices (CCDs) convert light to electrical charge and store the electrical charge for subsequent readout. In the case of amorphous silicon devices, a scintillating layer receives X-rays and generates light in proportion to the intensity of the received X-rays. An array of amorphous silicon photodiodes then converts and stores the generated light as electrical charge. For example, a photodiode of an amorphous silicon flat panel accumulates charge in proportion to an intensity of light received from an associated radiation source. After a specified time period, the accumulated charge is read in order to calculate the intensity of an image pixel associated with the photodiode. Accordingly, the accumulated charge is preferably directly proportional to the received light.
However, the photodiodes of the amorphous silicon sensors require a small bias voltage for proper operation. This bias voltage generates a small “dark current” that may cause a charge to accumulate within the photodiode that is unrelated to the intensity of the received light. This dark current thereby causes errors in the calculated intensity of the associated image pixel. Other imaging devices that convert radiation to electrical charge suffer from similar dark current problems.
Digital imaging devices as used in radiology or radiotherapy are typically flat panel imaging devices with amorphous silicon sensors deployed in a two-dimensional array. Such amorphous silicon (aSi:H) flat panel detectors are for example used as electronic portal imaging devices (EPID) on linear accelerators to a) image patient's anatomy or test phantoms and verify beam's aperture with high energy megavoltage (MV) beams delivered for treatment in radiotherapy, either in still planar views or in video mode (online) or in three-dimensional (3D) cone beam (CB) reconstructions, b) measure doserates or doses absolutely and relatively for machine quality assurance (maintenance, calibration of beam limiting and delivering devices) and in vivo (back projection to derive delivered dose distributions within a patient), or c) image patients' anatomy or test phantoms with kV beams delivered from an additional X-ray source to enable 2D, 3D and 4D (time-dependant) analysis.
In radiotherapy, aSi:H flat panels are widely used routinely for daily verification of patient's setup, and image quality and stability is of increasing relevance in the upcoming field of adapted image guided radiotherapy (IGRT) techniques. Typically, a MV panel on a standard linear accelerator (linac) has to capture images of several beams per patient every 10 to 15 min under very different conditions (anatomy, doserate, monitor units, beam energy, temperature). Since the device is irradiated with high energy photon beams, scattered dose (and occasionally the primary beam) damages the electronic components of the readout system (amplifiers) of the panel, so that the average life span of a clinically useable panel under such conditions may often not exceed 18 months.
Several approaches have been taken in an attempt to address the foregoing problems. One approach applies image processing techniques to each image frame that is produced from electrical charges read from an array of imaging elements. Known as offset correction, this approach involves acquiring image frames during a period of non-irradiation, calculating an average image frame from the acquired frames, and subtracting the average image frame from each frame acquired during subsequent radiation of the imaging elements. The averaged image frames are preferably acquired at the same rate as the subsequently-acquired frames so as to better approximate the effect of dark current on the subsequently-acquired frames. Since the extent of dark current effects varies across imaging devices, imaging devices are often sold with customized software for performing offset correction.
Furthermore, a linear relationship between doserate (or dose per frame) and pixel signal is conventionally assumed to achieve clinically useful images, which is roughly correct for newer panels. Although non-linear effects are known, the change of gain is relatively small at all higher doserates so that a fairly linear behaviour is assumed. Therefore, many sites in clinical and research environments found a sole background (or offset) and a linear gain correction to be sufficient, even if the panel was used for dosimetric purposes. A gain image describes a constant (linear) slope, which is assumed to be the pixel sensitivity. Conventional panel applications provide methods to acquire and store the gain image for a specific panel. For this purpose, the panel has to be irradiated with a flood field at a constant, higher doserate. The gain can simply be derived by dividing the doserate (normalized to a 16 bit value lower than or equal to hFFFF=65535) by the signal that has previously to be reduced by a provided offset value. Basically, this describes a two point measurement (background at doserate 0 and flood field at a high doserate), and considers linearity in between. Gain images were found to be relatively stable in time. For this reason, it is quite often decided to never recalibrate the gain during the life span of a panel.
However, offset correction, gain correction or recalibration often fails to provide suitable improvements with respect to image quality and life span of aged panels. Additional or alternative methods and systems are therefore desirable.